Ultrasonic diagnostic imaging with interpolated scanlines

ABSTRACT

An ultrasonic imaging method and apparatus are described for imaging the coronary arteries of the heart. The vascular system is infused with an ultrasonic contrast agent. A volumetric region of the heart wall including a coronary artery is three dimensionally scanned. A projection image of the volumetric region is produced from the scanning, providing a two dimensional contrast image of the coronary artery with the appearance of an angiogram. Preferably the coronary artery signals are segmented from contrast signals emanating from the myocardium and the heart blood pool so that the coronary arteries are clearly highlighted and distinct in the ultrasonic angiogram.

[0001] This invention relates to ultrasonic diagnostic imaging systemsand, in particular, to the use of ultrasonic diagnostic imaging systemsto image the coronary arteries.

[0002] Early detection of coronary artery disease is important for thetreatment and prevention of myocardial infarction, the primary cause ofdeath of adults in the world. One of the principal methods of detectionof coronary artery disease at present is the diagnostic angiogram. Anangiogram is acquired by injecting a radiopaque dye into the vascularsystem, usually by means of a catheter. The radiopaque dye infuses thecoronary arteries, and a radiological projection is made of the infusedarteries onto a radiographic sensor. The resultant angiogram will revealthe lumens of the arterial vessels of the heart as the radiopaque dyeflows through them. A narrowing of the infused lumen will provide anindication of an obstruction of a vessel and a potential condition forinfarction.

[0003] Ultrasound has been considered as a possible modality to use forcoronary artery examinations, which would have the advantage ofeliminating the exposure of the patient to the radiation used to formthe angiogram, to radiopaque dyes, and the surgical catheterizationprocedure. However, ultrasonic imaging has its own limitations. One isthat the major coronary arteries are located on the irregularly curvedsurface of the heart and traverse tortuously along the epicardialsurface of the heart. Thus, the coronary arteries cannot be viewed in asingle image plane, the most prevalent way ultrasonic imaging is done.Furthermore, imaging of the coronary arteries is impeded by the ribcage, which largely blocks ultrasound, and by the motion of the heartitself. Thus, even when a portion of the coronary arteries is accessibleto ultrasound, the images of the coronaries are likely to be fleeting,blurred, and of relatively poor resolution.

[0004] In accordance with the principles of the present invention, atechnique and apparatus are provided for ultrasonically imaging thecoronary arteries. The technique includes the use of an ultrasoniccontrast agent to sharply reveal the coronary arteries against theirbackground of the myocardium and the lungs, even when the heart is inmotion. The apparatus includes a three dimensional ultrasonic imagingsystem which is capable of scanning a three dimensional volume whichincludes at least a portion of the coronary arteries. In one displayformat the three dimensional volume including a coronary artery isprojected onto a display plane to produce an ultrasonic image in thesame manner as an angiogram. In another display format the threedimensional volume is displayed together with a realtime two dimensionalimage of a plane of the three dimensional volume. In another displayformat the bloodflow path of a coronary artery is displayed in aseparate, “straightened” rendering of the vessel.

[0005] Conventional radiological arteriograms only detect a few branchesdown the arterial tree from the major coronary trunks. Ultrasound iscapable of also visualizing transmural arteries. Thus, the presentinvention has the potential of providing visualization of vessels notseen with conventional angiography.

[0006] In the drawings:

[0007]FIG. 1 illustrates coronary arteries traversing the surface of aheart;

[0008]FIG. 2 illustrates an angiogram of a portion of the coronaryarteries shown in FIG. 1;

[0009]FIG. 3 illustrates a three dimensional volume intersecting aportion of the wall and chamber of a heart;

[0010]FIGS. 4a-4 c illustrate the sequential infusion of the chamber,myocardium and coronary arteries of the portion of the heart inside thevolume of FIG. 3;

[0011]FIGS. 5a and 5 b illustrate “slit-o-vision” scanning of a threedimensional volume with a linear and a phased array transducer;

[0012]FIGS. 6a and 6 b illustrate mechanical scanning of a threedimensional volume with a linear and a phased array transducer;

[0013]FIGS. 7a and 7 b illustrate electronic slit-o-vision scanning of athree dimensional volume with two dimensional linear and phased arraytransducers;

[0014]FIG. 8 illustrates an ultrasonic diagnostic imaging systemconstructed in accordance with the principles of the present invention;

[0015]FIG. 9 illustrates the partitioning of beamforming between ascanhead and an ultrasound system;

[0016]FIGS. 10a and 10 b illustrate the steering of a beam in theelevation direction by a scanhead beamformer;

[0017]FIGS. 11a, 11 b and 11 c illustrate different embodiments of ascanhead elevation beamformer;

[0018]FIG. 12 illustrates an azimuth beamformer which operates with theelevation beamformers of FIGS. 11a, 11 b, and 11 c;

[0019]FIG. 13 is a plan view of a two dimensional transducer array forthree dimensional scanning in accordance with the present invention;

[0020]FIG. 14 illustrates a receive sub-aperture of the transducer arrayof FIG. 13;

[0021]FIGS. 15a-15 g illustrate different transmit sub-apertures of thetransducer array of FIG. 13;

[0022]FIG. 16 illustrates scanhead microcircuitry for sampling thesignals received by a transducer element of the transducer array of FIG.13 in a desired time relationship;

[0023]FIG. 17 is a more detailed view of the microcircuitry of FIG. 16;

[0024]FIG. 18a illustrates a scanhead micro-beamformer and multilinebeamformer system suitable for processing the signals received by thetransducer array of FIG. 13;

[0025]FIG. 19a illustrates operation of the system of FIG. 18a for ahexagonal scanning pattern;

[0026]FIGS. 19b and 19 c illustrate the use of interpolation to developa hexagonal scanline pattern;

[0027]FIG. 18b illustrates the use of a multiline scanheadmicro-beamformer in combination with a system multiline beamformer;

[0028]FIGS. 18c and 18 d illustrate single line and multiline beamsteering from a 2D transducer array patch;

[0029]FIG. 20 illustrates a three dimensional volume containing a twodimensional image plane;

[0030]FIG. 21 illustrates the time interleaved sampling of the threedimensional volume and two dimensional image plane of FIG. 20;

[0031]FIG. 22 illustrates a duplex display of the three dimensionalvolume and two dimensional image plane of FIG. 20;

[0032]FIG. 23 illustrates a duplex display of a large three dimensionalvolume and a smaller three dimensional volume contained within thelarger volume;

[0033]FIG. 24 illustrates a three dimension image volume containingcoronary arteries;

[0034]FIG. 25 illustrates an algorithm for detecting the center of ablood vessel in a three dimensional image; and

[0035]FIG. 26 illustrates a “straightened” display of one of thecoronary arteries of FIG. 23.

[0036] Referring first to FIG. 1, a picture of a heart 10 is shown.Located on the outer surface of the heart are the coronary arteries 12,which provide a continuous supply of blood to the heart muscle, themyocardium. The outer surface of the heart is irregularly rounded withperiodic depressions and elevations, and the coronary arteries locatedon this surface follows this continuously bending surface and its highand low points. Thus, the coronary arteries are not located on a planarsurface, but on a surface which undergoes many curves and contortions.The coronary arteries cannot be imaged by a single plane which bisectsthe heart, but by techniques which will image the three dimensionalpaths of the coronary arteries and all of their bends, twists, andturns.

[0037]FIG. 2 depicts an angiogram of the coronary arteries of the heartof FIG. 1. The angiogram of FIG. 2 is formed by first injecting aradiopaque dye into the body which infuses the coronary arteries. Abroad beam of x-rays is then transmitted through the chest of thepatient and onto a radiographic plate on the opposite side of thepatient. The radiographic plate is continually scanned to create animage. The x-rays which pass through the heart without intersecting thecoronary arteries will appear as bright areas in the image, but x-rayswhich strike a dye-infused artery will not reach the radiographic plate,leaving an x-ray “shadow” 14 of the coronary arteries on the plate. Theresulting shadow image of the coronary arteries will appear as shown inFIG. 2. Since a large area or even the full heart is illuminated withx-rays, the twisting and turning coronary arteries on the heart surfacewill leave a pattern in the transmitted x-rays, even though their twistsand turns extend in three dimensions. Obstructions in the coronaryarteries will be revealed by sudden changes in the width and/orbrightness of an arterial “shadow.”

[0038]FIG. 3 depicts the ventricular region of a heart. The leftventricle LV and right ventricle RV are depicted on the drawing. Aportion of the heart wall and left ventricle are to be imaged in thisexample, and are contained within an imaging volume 20. The chambervolume of the left ventricle is indicated at 18, the myocardium isindicated at 16, and the coronary arteries 12 are located on the outersurface of the myocardium. An ultrasonic contrast agent is introducedinto the body of the patient and the imaging volume is scannedultrasonically for both the harmonic and the nonlinear fundamentalreturn from the contrast agent. Initially, before any of the contrastagent has reached the heart, there will be no harmonic return exceptthat caused by nonlinear propagation of the ultrasonic signal, whichwill be relatively low in intensity. This tissue harmonic return can bereduced by pre-distorting the transmitted pulse as described in U.S.Pat. No. 5,980,457 or by thresholding or other techniques. When thecontrast agent reaches the heart through the circulatory system it willinitially fill the chamber of the left ventricle in the imaging volume20, as depicted in FIG. 4a. The left ventricle will light up with astrong harmonic return and will appear brightly in the ultrasonic image.As the contrast agent is pumped from the heart it will next infuse thecoronary arteries 12, as depicted in FIG. 4b. At this stage both theheart chamber 18 and the coronary arteries 12 will appear brightly inthe ultrasonic image. Finally the contrast agent will perfuse thecapillary bed of the myocardium 16 from the coronary arteries 12. Thethree regions 12, 16 and 18 will then appear brightly illuminated by thecontrast agent return echoes as depicted in FIG. 4c.

[0039] Since the purpose of the procedure is to examine the coronaryarteries with as little clutter from other tissue as possible, it is thesecond stage of infusion, that depicted in FIG. 4b, which is of primaryinterest. Hence the clinician should be recording the sequence of eventsso as to capture the ultrasonic images when the coronary arteries areinfused with contrast agent and before the myocardium becomes perfused,since harmonic signals from the myocardium are unwanted and would beregarded as clutter. The harmonic return from the LV chamber is alsoconsidered clutter as it is unwanted harmonic signals and can interferewith the projection of the infused coronary arteries onto an imageplane. When images such as those depicted in FIG. 4b have been captured,they are preferably processed to eliminate the unwanted signal from theheart chamber 18. This can be done by adaptive beamforming and/or byimage post-processing which spatially mask out the unwanted imagesignals. Since the left ventricle 18 is infused first, an algorithmwhich detects the initial harmonic return and then masks out image areascontiguous with the initial harmonic return within one or two heartcycles can effectively remove heart chamber signals from the image.Another technique for masking the left ventricular return is torecognize that the left ventricle is a large blood pool and eliminatedisplaying signals in large area which lack tissue return signals.Thresholding can be used to segment between blood pool signals whichhave a blood signal density of 100% and the approximately 6% bloodsignal density of tissue. Another sensing technique is to operate in thefundamental frequency spectrum and recognize that fundamental frequencysignals returned from blood are of a lower amplitude than echoesreturned from tissue. Harmonic signals returned from areas withpredominately low amplitude fundamental signal returns are eliminatedfrom the display to mask the left ventricle. Yet another technique is torecognize that bloodflow velocity in the left ventricle is greater thanthat in the coronary arteries or myocardium, and mask out the highestvelocity signals from the display. The remaining bright image signalsfrom the coronary arteries can then be displayed in a three dimensionaldisplay, or projected onto a darkened image plane as by threedimensional maximum intensity rendering to produce a two dimensionalultrasonic projection image of the coronary arteries which will appearmuch like an angiogram and hence can be used for diagnosis by cliniciansfamiliar with angiograms.

[0040] The segmentation of image areas into blood and myocardium canalso be used to adaptively beamform in order to increase spatialresolution and temporal resolution in the myocardium. Such a processwould generally use larger transmit apertures (with less beam width butrequiring more transmit cycles) in the regions of interest (e.g., themyocardium). Smaller transmit apertures are used in the blood poolregions to permit higher-order multiline, and thereby provide fasteracquisition. In addition to these aperture adaptations, adaptivebeamforming techniques which enhance the imaging of the coronaryarteries include sensing the intensity of received signals so as todetect the high intensity return signals from the contrast agent infusedblood pool of the left ventricle, and responding by reducing transmitpower to the left ventricle, since the left ventricle signals areunwanted and reduced transmit power will cause less disruption of theagent in the left ventricle which is being pumped to the coronaryarteries. Another alternative when using multiline reception, whichrequires a “fat” (broad) beam to insonify multiple scanlinessimultaneously, is to adaptively alter the aperture and hence narrow thetransmit beam profile when overlapping the left ventricle, since thebeam need only be broad enough to insonify the coronary arteries and notthe coronary arteries and the left ventricle. Yet another adaptation isto optimize time-of-flight adjustments for tissue and not blood, and toinhibit these adjustments when receiving blood pool signals from theleft ventricle, which is not the target of interest. Yet a furtheradaptation is to adaptively tailor the aperture in consideration oftransducer elements which are blocked by the ribs when imaging the hearttransthoracically.

[0041] It has been found that the coronary imaging procedure can beconducted to prevent the third stage shown in FIG. 4c from effectivelyoccurring, thereby sustaining the desired second stage of infusion shownin FIG. 4b where the coronary arteries are clearly segmented. The finecapillary bed of the myocardium 16 is perfused slowly by a very smallnumber of microbubbles of the contrast agent. The velocity of bloodflowin a coronary artery supplying blood to a capillary bed will berelatively high, and as the volume of blood is distributed over the manycapillaries of the bed the velocity becomes considerably lower in thecapillaries. The bloodflow velocity increases again in the largecollector vessels at the output of the capillary bed. Approximately fiveto eight cardiac cycles can be required to initially perfuse a capillarybed of the myocardium, or to reperfuse the capillaries after thecontrast agent has been disrupted. By balancing the image frame rate(i.e., the transmit pulse rate) and the transmit pulse power, these verysmall microbubbles can be continuously disrupted as they begin toperfuse the myocardium. A higher frame rate and/or a higher transmitpower will cause increased microbubble disruption in selected areas ofthe cardiovascular system. It has been found that transmitting with amechanical index setting of approximately 0.1 or less will cause littleor no microbubble disruption, enabling the blood pool of the heartchamber, coronary arteries, and the myocardial capillary bed to beimaged at these levels. Above this level noticeable microbubbledisruption will occur. In the range of a mechanical index ofapproximately 0.2-0.5, microbubbles are disrupted in substantial numbersbefore they are able to reperfuse the capillary bed in the image plane.Thus, imaging at these power levels will result in significant harmonicsignal returns from the heart chamber and coronary arteries, which aresignificantly reinfused in the image plane between transmit pulses, withlittle signal returns from the capillary bed. At higher power levels,and particularly above a mechanical index of 1.0, there will besubstantial microbubble disruption in the heart chamber, coronaryarteries, and capillary bed, with only minor reinfusion of the imageplane in the heart chamber blood pool. The exact numbers will vary forparticular contrast agents. By use of a suitable low but disruptivemechanical index setting, the capillary bed of the myocardium can beeffectively kept free of substantial amounts of contrast agent and hencewill produce little if any harmonic contrast return signals. The resultis that the second stage of FIG. 4b, where only the heart chamber andcoronary arteries are producing significant harmonic return signals, canbe maintained for a considerable period of time by proper selection offrame rate and transmit power. Combining selective disruption of thecapillary bed with masking of the left ventricle blood pool enable thecoronary arteries to be segmented for display. Another variation is torecognize the low velocity of flow in the myocardium as noted above, andto mask or reject signals of low flow velocity from the capillary bed.

[0042] Furthermore, the coronary artery imaging procedure can becontrolled to reduce unwanted harmonic returns from contrast agent inthe heart chamber, thereby minimizing the need for masking or other echoelimination techniques. The contrast agent can be administered throughan intravascular catheter, which can be threaded into the aortic root.When the contrast agent is injected at the aortic root, the flow ofblood out of the heart will prevent immediate entry of contrast agentinto the heart chambers. Since the coronary arteries receive their bloodsupply from the aortic root, the injection of contrast agent at thislocation will cause the coronary arteries to immediately become infusedwith the agent. Thus the coronary arteries will be the first structureto light up with contrast agent, which will not enter the heart untiltraversing the vascular system and returning to the heart by venousflow, at which point a significant amount of agent may be eliminated bylung filtering. Thus, clutter from contrast agent in the heart chambersis reduced if not eliminated for at least the initial period of theprocedure.

[0043] One way to form a planar ultrasonic projection image of avolumetric region is by means of the technique known as “slit-o-vision,”which is described in U.S. Pat. No. 5,305,756. In the slit-o-visiontechnique a volumetric region is insonified with an ultrasonic beamwhich is divergent in the elevation dimension and focused in the azimuthdimension. The technique takes its name from the fact that such an imagecan be formed by use of an aperture which is relatively long in theazimuth dimension and narrow in the elevation dimension. An ultrasoundbeam produced from such an aperture utilizes diffraction to radiate anessentially cylindrical wavefront that, while focused in azimuth,develops the desired divergence in the elevation direction. Anelevationally divergent beam can also be produced by acoustic lenses orelectronic lenses. For instance, when an element of an array is convexin the elevation dimension or has a lens which is elevationallydivergent, elevationally divergent beams can be produced by a lineararray 10 as shown in FIG. 5a. The elevation and azimuth dimensions areindicated by the EL and AZ arrows, respectively. The elevationallydivergent beams will insonify a wedge-shaped volume 30. Points which areof the same range locus from the array 10, such as those along rangelocus 28, are acoustically integrated and projected at that range ontoprojection plane 32. This acoustic integration and projection occurs atevery range in the volume, so that the entire volumetric region 30 isprojected onto the projection plane 32. If the volumetric regioncontains only or principally signals from coronary arteries, thecoronary arteries in the volume will be projected onto the projectionplane 32 and appear as the image. FIG. 5b illustrates the same resultsfrom use of a phased array transducer 10′, in which case the volumetricwedge 60 is more triangular and projects onto a triangular plane 62.

[0044] In FIGS. 6a and 6 b the same volumetric regions 30 and 60 areinsonified by sweeping beams focused in elevation and azimuth over thevolumetric regions. This is done by rocking the 1D array in theelevation direction as shown by the arrows 42. After image planes withinthe volumes 30,60 have been acquired and contrast signals from thecavity have been removed, points of common range such as those alongconstant range locus 28 are integrated together to project the planesonto a common projection plane 32,62. Once again, an image in the formof an angiogram is formed when coronary artery information ispredominate in the volumetric regions 30,60.

[0045]FIGS. 7a and 7 b illustrate another slit-o-vision technique, inthis case by the electronic synthesis of elevationally divergent beams,which has better sensitivity but generally lower frame rate than theembodiment of FIG. 5. These embodiments use two dimensional arrays 10″,which may be operated as either 1.5D or 2D arrays. The transmit beamsare made elevationally divergent by pulsing the central elevationelements first, then proceeding to pulse the outermost elevationelements last. FIG. 7a depicts a linear array scanning format whichscans volumetric region 30, and FIG. 7b depicts a phased array scanningformat which steers beams in a volumetric region 60. As in the case ofthe FIG. 5a and 5 b embodiments, points at a common range locus in thevolumetric regions acoustically integrate onto a projection plane 32 or62 for a projection image of the structure in the volumetric regions 30and 60. Alternatively, the 2D array of FIG. 7 can be focused and steeredin elevation to effect the scanning of the volume as shown in FIG. 6.

[0046] Where discrete planes of a volumetric region are acquired asdescribed below, they can be combined to form a projection image asshown in FIG. 2 using volume rendering techniques. This threedimensional method has lower frame rates than the above “slit-o-vision”approaches, but provides greater spatial resolution, less clutter fromoff-axis reflectors, and more control over the image by variation of therendering parameters.

[0047]FIG. 8 illustrates an ultrasonic diagnostic imaging systemconstructed in accordance with the principles of the present invention.A scanhead 26 including an array transducer 10 is connected by a cable11 to a beamformer 36. The beamformer controls the timing of actuationsignals applied to the elements of the transducer array for thetransmission of steered and focused transmit beams, and appropriatelydelays and combines signals received from the transducer elements toform coherent echo signals along the scanlines delineated by thetransmit beams. The timing of the beamformer transmission is alsoresponsive to an ECG signal when it is desired to synchronize or gateimage acquisition with a particular phase of the heart cycle. Thebeamformer is further responsive to a scanhead position signal when thetransducer is being mechanically moved to sweep ultrasonic beams over avolumetric region, thereby enabling beams to be transmitted when thetransducer is properly oriented with respect to the volumetric region.

[0048] The output of the beamformer is coupled to a pulse inversionprocessor 38 for the separation of fundamental and harmonic frequencysignals. Pulse inversion processors are well known in the art and aredescribed in U.S. Pat. Nos. 5,706,819 and 5,951,478. These patentsdescribe how echoes from alternately phased pulses can be used toseparate harmonic contrast signals from fundamental signals, which is apreferred method of separating signals from contrast agents for coronaryimaging in accordance with the present invention.

[0049] The fundamental and/or harmonic signals may be B mode processedor Doppler processed, depending upon the desired information to bedisplayed. For Doppler processing the signals are coupled to a wallfilter 22 which can distinguish between flow, stationary tissue, andmoving tissue. A preferred wall filter for contrast imaging is describedin U.S. patent [application Ser. No. 09/156,097], which is also capableof performing harmonic contrast signal separation. The filtered signalsare applied to a Doppler processor 42, which produces Doppler power,velocity, or variance estimation. A preferred Doppler processor forharmonic Doppler signal estimation is described in U.S. Pat. No.6,036,643. Artifacts from scanhead motion which can contaminate Dopplerimaging are removed by a flash suppressor 44. Various techniques may beused to remove flash artifacts prior to or subsequent to imageformation, including the notch filter technique described in U.S. Pat.No. 5,197,477 and the min-max filter technique described in U.S. Pat.No. 5,782,769. The processed Doppler signals are stored in a Dopplerimage memory 40′.

[0050] Signals which are to be B mode processed are applied to a B modeprocessor 24 which detects the signal amplitude. B mode processedsignals are stored in a tissue image memory 40.

[0051] The B mode and Doppler signals are applied to a coordinatetransformation processor 46. For conventional two dimensional imagingthe coordinate transformation processor will function as a scanconverter, converting polar coordinates to Cartesian coordinates asnecessary and filling spaces between received lines with interpolatedimage data. The scan converted images are coupled to a video processor70 which puts the image information into a video format for display ofthe images on a display 100. The images are also coupled to a Cineloop®memory 56 for storage in a loop if that function is invoked by the user.

[0052] When 3D imaging is being performed by the ultrasound system, thecoordinate transformation processor may be used to scan convert thetissue and Doppler signals in planes of image information over thescanned volume, or may be used to transform the coordinates of the imagedata into a three dimensional data matrix. Preferably the coordinatetransformation processor operates in cooperation with a volume renderingprocessor 50, which can render a three dimensional presentation of theimage data which has be processed by the coordinate transformationprocessor. Three dimensional images of tissue are rendered in accordancewith tissue rendering parameters 54 which are selected by the userthrough a control panel or user interface (UIF). Three dimensionalimages of Doppler information are rendered in accordance with blood flowrendering parameters 52. These parameters control aspects of therendering process such as the degree of transparency of tissue in thethree dimensional image, so that the viewer can see the vasculatureinside the tissue. This capability is important when 3D images of bothtissue and flow are being rendered, as described in U.S. Pat. No.5,720,291. Three dimensional images can be stored in the Cineloop®memory 56 and replayed to display the scanned volume in a dynamicparallax presentation, for instance. A three dimensional rendering offlow without the surrounding tissue, as described in U.S. Pat. Re36,564, can reveal the continuity of flow of blood vessels andobstructions in those vessels and is useful for coronary arterydiagnosis in accordance with the present invention.

[0053] Different transducer probes can be used to scan a volumetricregion of the heart which includes the coronary arteries. Either a 1D(azimuth steered) or a 1.5D or 1.75D (azimuth steered and elevationfocused) array may be moved mechanically to sweep beams over the threedimensional volume. For electronic steering either a 1.75D (minimallyelectronically steered in azimuth and elevation) or a 2D (fullyelectronically steered in azimuth and elevation) array may be used. Anembodiment which uses a 2D array transducer 10″ is shown in FIG. 9. Animportant consideration in the use of two dimensional arrays is thenumber of cable wires used to connect the probe to the ultrasoundsystem. Various approaches can be used to reduce the number of cableconductors and thus the size of the cable, including wireless links tothe ultrasound system, micro-beamforming in the probe, digital or analogtime multiplexing, the use of sparse arrays, and the use oftransmit/receive multiplexers. One solution is an r.f. probe whichtransmits echo signals wirelessly to the ultrasound system as describedin U.S. patent [application Ser. No. 09/197,398]. Another solution, whena cable connection is used, is to partition the beamformer between thescanhead and the ultrasound system as described in U.S. patent[application Ser. NO. 09/197,196]. The embodiment of FIG. 9 makes use ofthis approach by performing elevation beamforming in the scanhead 26 andazimuth beamforming in the ultrasound system 101. For example, supposethat the two dimensional array has 128 columns of elements extending inthe azimuth direction (indicated by the AZ arrow in the drawing) and sixrows of elements in the elevation direction (indicated by the EL arrow).If each element of the array were connected by its own conductor to theultrasound system, a cable of 768 signal conductors would be required.In the embodiment of FIG. 9 each column of six elements is coupled to anelevation beamformer 36 a which appropriately excites (on transmit) anddelays and combines (on receive) signals from the six elements of thecolumn. This combines the six signals in each column into one elevationbeamformed signal, which is then coupled over a cable conductor to theultrasound system, where the elevation beamformed signals are beamformedin the azimuth direction. In the foregoing example, the 128 elevationbeamformed signals are coupled over the 128 conductors of a cable 11, asignificant reduction in cable size as compared to a probe withoutscanhead beamforming. At least elevation steering is performed in theelevation beamformer 36 a, and preferably both steering and focusing areperformed in the elevation beamformer.

[0054] The operation of the elevation beamformer is illustrated in FIGS.10a and 10 b. In FIG. 10a a beam is being steered normal to the arraytransducer as indicated by the 0° arrow extending from the elements 10 ₁through 10 _(n), which comprise a column of elements in the elevationdirection. Signals at the center of the column are delayed more thansignals at the ends of the column as indicated by the relative length ofthe delays 102 for the different elements to effect a focus. Delayedreceive signals are combined by a summer 104, then coupled over a signallead in the cable 11 to the azimuth beamformer 36 b. FIG. 10billustrates the situation when a beam is to be transmitted or receivedfrom the left at a 30° inclination in elevation as indicated by the 30°arrow. In this case signals on the left side of the array are moregreatly delayed as indicated by the relative length of the delays 102.Received signals are combined by the summer 104 and coupled through thecable to the azimuth beamformer 36 b.

[0055]FIGS. 11a-11 c illustrate the implementation of the elevationbeamformer in three different ways (neglecting any buffering or gainelements). FIG. 11a illustrates an analog implementation in which eachtransducer element 10 _(m) is coupled to an analog delay line 106. Thelength of the delay is set by choosing the input or output tap of thedelay line and the delayed signals are coupled to an analog summer or toan A/D converter if the signals are to be digitally combined. In FIG.11b each transducer element 10 _(m) is coupled to a CCD delay line 108.The length of the delay is set by choosing an input or output tap thatdetermines the number of charge storage elements in the delay line or byvarying the rate at which the charge samples are passed through thecharge storage elements. The outputs of the delay lines are summedeither in sampled analog format or after being digitized.

[0056]FIG. 11c illustrates a digital embodiment of an elevationbeamformer. In this example the elevation beamformer has 128sub-beamformers 120, each processing the signals from one elevationcolumn of six transducer elements. Each of the transducer elements 101through lon is coupled to an A/D converter 110 and the digitized signalsare delayed by a digital delay line 112, which may be formed by a shiftregister, FIFO register, or random access memory. The appropriatelydelayed signals are combined in a summer 104 and coupled over cableconductors to the azimuth beamformer. To conserve cable conductors whenusing multibit signal samples, the data values from several of thebeamformer channels 120 can be interleaved (time multiplexed) and sentover the same group of conductors at a data rate sufficient for thedesired level of realtime imaging performance.

[0057]FIG. 12 illustrates the organization and control of a number ofbeamformer channels 120 of a scanhead elevation beamformer. Thebeamformer comprises N elevation sub-beamformers 120 ₁-120 _(n) whereeach sub-beamformer receives signals from a column of transducerelements in the elevation direction, as indicated by the number 6 forthis example. Data to control the elevation beamforming (such aselevation angle and focusing) is sent to a timing & delay decoder & datastore 126 in the scanhead 26, preferably serially over a cableconductor. This control data is decoded and delay values coupled to adelay control 124, which sets the beamformer channels for the desireddelays for each transducer element. For dynamic focusing the delays arechanged as echoes are received. The elevation aperture can be varied byapplying zero weights to some of the outermost channels when a smaller(near field) aperture is desired. The data received by the timing &delay decoder & data store 126 is also used to control transmit timingby pulse transmitters 122 ₁-122 _(n), each of which controls thetransmission of the six transducer elements in an elevation column inthis example. When received echo signals are processed in the analogdomain as illustrated by FIGS. 11a and 11 b, the signals from the 128channels of the elevation beamformer in this example are sent over 128cable conductors to the azimuth beamformer 36 b. When the echo signalsare processed digitally the signals from the 128 channels areinterleaved (time multiplexed) and sent over digital conductors of thecable 11 to the azimuth beamformer in the ultrasound system 101.

[0058] A true 2D electronically steered embodiment of the presentinvention is illustrated starting with FIG. 13. This drawing shows aplan view of a 2D transducer array 200 of greater than three thousandtransducer elements. For ease of illustration the small boxes in thedrawing which represent individual transducer elements are shown spacedapart from each other. However, in a constructed embodiment, theindividual transducer elements are close packed in a repeating hexagonalpattern. The 2D array has an overall dodecahedral outline. In apreferred mode of operation beams are transmitted outward from thecenter of the array and can be steered and focused in a cone of at least±30° about a line normal to the center of the array. When steeredstraight ahead, echoes received from along a transmitted scanline areinitially received at the center of the array and then in circular orarcuate groupings of elements centered on and extending outward alongthe projection of the scanline onto the surface of the array. In theillustrated embodiment approximately the central one-quarter of theelements are used for beam transmission. The entire array is availablefor echo reception.

[0059] The array 200 of FIG. 13 is seen to be drawn in alternate lightand dark groupings 202 of twelve transducer elements. One of thesegroupings 202, referred to herein as a “patch” of transducer elements,is shown in a separate enlarged view in FIG. 14. These irregularhexagonal patches 202 of twelve elements are beamformed together duringecho reception as discussed in detail below. Elements in the center ofthe array (approximately 750 elements) are connected in groups of threefor transmission by high voltage mux switches. FIGS. 15a-15 f show someof the three-element configurations that are possible during beamtransmission. The transmit groupings can also simply be three elementsadjacent to each other in a straight line. The exact configuration orconfigurations used to transmit a given beam depend upon the desiredbeam characteristics and its azimuth. Four elements may also beconnected together for transmission as illustrated by the diamond shapedgrouping of four elements in FIG. 15g.

[0060] Since a cable with more than three thousand conductors is notcurrently practical, each patch of twelve elements of the array isbeamformed in the scanhead. This reduces the number of signals whichmust be coupled to the ultrasound system beamformer to approximately256. Then, a 256 channel beamformer in the ultrasound system can be usedto beamform the partially beamformed signals from the scanhead.

[0061] Because the elements of each receive patch of twelve elements ofthe 2D array are sufficiently small, contiguously located, and closelypacked, the echo signals received by the elements of a patch will bealigned to within one wavelength at the nominal receive frequency forsteering angles of approximately 40° or less (neglecting focal delays).The echoes of the elements are then sampled to bring all of the patchelement signals into precise time alignment. The sampling is done with arange of sampling delays with a precision of a fraction of a wavelengthto bring the signals from all of the patch elements to a time alignmentwithin the precision of the sampling clock quanta, preferably {fraction(1/16)} of a wavelength or less. The time-aligned signals from the patchelements are then combined. This beamforming of each patch is done bymicroelectronics located immediately behind the transducer array in thescanhead to facilitate interconnections. Sample time shifting andalignment is performed by the sampling delay line shown in FIGS. 16 and17. Each element 204 of a patch of elements which is to be partiallybeamformed is coupled by way of an amplifier 206 to a sampling inputswitch 208. The sampling input switch 208 is continually conductingsamples of the transducer signal onto capacitors 212 in a sequentialmanner. The sequencing of the switch 208 is under control of a ringcounter 210 which is incremented by a clock signal. As the darkenedsegment of the ring illustrates, the sampling input switch iscontinually sampling the input signal onto successive ones of thecapacitors 212 in a circular manner. The amplifier 206 has a bipolaroutput drive so that the charge of a capacitor can be either increasedor decreased (discharged) to the instantaneous signal level at the timeof sampling.

[0062] The signal samples stored on the capacitors 212 are sampled by asampling output switch 214 which samples the stored signals in asequential manner under control of a second ring counter 216. As shownby the darkened segment on the ring of the second ring counter 216, thesampling output switch 214 samples the stored signals in a particulartime relationship to the input switch and its ring counter. The timedelay between the input and output sampling is set by a time shifter 220which establishes the time delay between the two ring counters. Thus thetime of sampling of the output signal samples can be incrementallyadvanced or delayed as a function of the timing difference between thetwo ring counters. This operation can be used to bring the output signalsamples of all the elements of a patch into a desired time alignmentsuch as the sampling time of a central element of the patch. When thesignals from all of the elements of the patch are within a desired rangeof sampling time, the signals can be combined into one signal forfurther beamforming in the ultrasound system. The time aligned outputsignals are further amplified by an amplifier 218 and coupled to asummer for combining with the signals of the other elements of thepatch.

[0063] Details of a constructed embodiment of the arrangement of FIG. 16are shown in FIG. 17. In integrated circuit fabrication the samplingswitches do not have rotating wipers as illustratively shown in FIG. 16,but are formed by a plurality of gates 228. Each of the gates 228 iscontrolled by the output of an output stage of a shift register 230,which is arranged to circulate one bit and thereby operate as a ringcounter. When the bit is shifted to a particular stage of the shiftregister 230, the gate 228 connected to that stage is closed to conducta signal sample to its capacitor 212. The output switches are similarlyconstructed as a series of parallel gates 234, and are similarlycontrolled by stages of circulating shift register 232. Signal samplestaken from the capacitors 212 are amplified and resistively coupled to acurrent summing node for summation with the other signals of thegrouping.

[0064] A clock command memory 240 is located in the scanhead andpreferably on the same integrated circuit as the sampling circuitry. Theclock command memory stores data identifying the time delays needed forone or more receive echo sequences. The control data for the currentbeam is coupled to a clock delay controller 242 which controls therelative time relationship between the two ring counters. The controller242 does this by blocking clock cycles applied to the first ring counter230 from reaching the second ring counter 232, or by insertingadditional clock cycles into the clock signal. By blocking or insertingshift register clock pulses to the second ring counter the relativetiming between the two ring counters is adjustably advanced or retarded.The time aligned samples from all of the transducer elements of thepatch are then combined at a current summing node I Node. The summedsignals from the patch are coupled through the scanhead cable to theultrasound system beamformer.

[0065] With the addition of a second sampling output switch for eachelement controlled in a different time relationship than the firstsampling output switch, and a second summer for the second samplingoutput switches of the patch elements, a second, receive beam can beproduced at the same time as the first receive beam. Thus, each patchbecomes a small multiline receiver receiving two (or more) receive beamssimultaneously, which is useful in the multiline embodiment describedbelow.

[0066] The microbeamformer for the patches can utilize otherarchitectures such as charge coupled delay lines, mixers, and/or tappedanalog delay lines.

[0067] Three dimensional imaging requires that the volumetric region besufficiently sampled with ultrasound beams over the entire volume. Thisrequires a great many transmit-receive cycles which causes the timeneeded to acquire a full set of volumetric data to be substantial. Theconsequences of this substantial acquisition time are that the framerate of a realtime 3D display will be low and that the images will besubject to motion artifacts. Hence it is desirable to minimize the timerequired to acquire the necessary scanlines of the volumetric region. Apreferred approach to this dilemma is to employ multiline beamforming,scanline interpolation, or both, as shown in FIGS. 18 and 19. Whilebeams may be steered in a square or rectangular pattern (when viewed incross-section) to sample the volume being imaged, in a preferredembodiment the beams are oriented in triangular or hexagonal patterns inthe volumetric region to sufficiently and uniformly spatially sample theregion being imaged. FIG. 19a is a cross-sectional view through thevolumetric region in which scanlines in the volumetric region areaxially viewed. In this example nineteen scanlines are produced forevery transmit beam. The scanline locations are spatially arranged inhexagonal patterns. The nineteen scanline locations of one hexagonalpattern are denoted by circles which represent axial views along thescanlines. The nineteen scanline locations are insonified by a “fat”transmit beam of a desired minimum intensity across the beam. Thetransmit beam in this example is centered on the location of scanline270, and maintains the desired acoustic intensity out to a peripherydenoted by the dashed circle 250, which is seen to encompass allnineteen scanline locations. The echoes received by the elements of thetransducer array are partially beamformed by a micro-beamformer 280 inthe scanhead as described above and coupled to a 19× multilinebeamformer 282 in the ultrasound system as shown in FIG. 18a. In thisexample a 2D transducer array of 3072 elements is operated in patches of12 elements, producing 256 patch signals which are coupled to theultrasound system by a cable 281 with 256 signal conductors withoutmultiplexing. The 19× multiline beamformer processes the 256 echosignals received from the transducer patches with nineteen sets ofdelays and summers to simultaneously form the nineteen receive scanlines252-274 shown in FIG. 19a. The nineteen scanlines are coupled to animage processor 284, which performs some or all of the harmonicseparation, B mode, Doppler, and volume rendering functions previouslydescribed in FIG. 8. The three dimensional image is then displayed onthe display 100.

[0068] Interpolation may be used to form scanline data, eitheralternatively to or in conjunction with multiline scanline formation.FIG. 19b illustrates a series of scanlines 361-367 marked by thedarkened circles which have been acquired from a volume being imaged ina hexagonal pattern as indicated by the background grid pattern. Thescanlines 361-367 can be acquired individually or in groups of two ormore by multiline acquisition. Scanlines at the undarkened circlelocations are interpolated from the acquired scanlines using twopointr.f. interpolation. The interpolated scanline 371 is interpolated byweighting each of the adjacent scanlines 361 and 362 by ½, thencombining the results. The weights used are a function of the locationof the scanline being produced in relation to the locations of the threereceived scanlines whose values are being interpolated. Similarly,interpolated scanline 372 is interpolated using adjacent scanlines 362and 367, and interpolated scanline 373 is interpolated using adjacentscanlines 361 and 367. Each group of three scanlines is used tointerpolate three intermediate scanlines using weighting factor whichare a factor of two (2⁻¹), enabling the interpolation to be performedrapidly by shifting and adding the bits of the data being interpolated.This avoids the use of multipliers and multiplication and affordshigh-speed processing advantageous for realtime 3D display rates.

[0069]FIG. 19c illustrates a further iteration of the interpolation ofFIG. 19b in which the scanline density of the volume is increased evenfurther by interpolation. In this illustration two further sets ofscanlines 381-383 and 387-392 are interpolated between the previous set.These scanlines may be interpolated using the previously interpolatedset of scanlines, or they may be interpolated directly (andsimultaneously, if desired) from the acquired scanlines 361,362,367.These scanlines also have the advantage of being weighted by weightingfactors which are a factor of two. The set of interpolated scanlinesmost central to the three received scanlines, 381-383, are interpolatedusing weighting factors of ½ and ¼. Scanline 381, for instance, isproduced by (½(scanline 361)+¼(scanline 362)+¼(scanline 367)). The outerset of scanlines is produced by ¼, ¾ weights as described in U.S. Pat.No. 5,940,123. Scanline 392, for instance, is produced by (¼(scanline367)+¾(scanline 361)) or, to avoid multiplication, (¼(scanline367)+¼(scanline 361)+¼(scanline 361)+¼(scanline 361)). FIG. 19cillustrates corresponding sets of interpolated scanlines for receivedscanlines 362,363,367, including the central group of scanlines 384-386,and the outer set of scanlines 393-396. To reduce motional artifacts,the received scanline data can be filtered in either r.f. or detectedform prior to display.

[0070] The above example uses a linear interpolation filter kernel. Itis also possible to use an interpolation kernel that has a non-linershape (such as, for example, cosine, sinc, etc.) However the filtercoefficients of these other filters will generally not have thedesirable power of two property.

[0071] The use of patches to reduce the size of the cable needed toconnect the scanhead to the ultrasound system may, under certainoperating conditions, give rise to undesired grating lobes in thescanhead's beam pattern. This is due to the grouping of individualtransducer elements into a single unit, giving the transducer array acoarser pitch, even with the use of micro-beamforming as describedabove. This problem can be reduced by considering each patch to be asub-aperture of the entire 2D array which is capable of receivingsignals from multiple, closely spaced scanlines in the transmit beamfield. The signals from the sub-apertures can be delayed and summed toform a group of multiline received scanlines. Grating lobes which ariseby reason of the periodicity of the sub-apertures and can contributeclutter to the final image are reduced by producing two or moredifferently steered signals from each sub-aperture (patch). The steeringdifference is kept small, within the beamwidth of the patch. By keepingthe steering delay profile less than {fraction (λ/2)}, significantgrating lobes are kept out of the image field.

[0072] A simple 1D example illustrates these effects. Consider asixty-four element 1D linear array with interelement spacing (pitch) of{fraction (λ/2)}. The array is divided into four patches of sixteenelements each. Two beams are steered to the left and right of a nominaldirection on each patch. The steering angles are limited so that otherlines or samples can be interpolated between these two receivedmultilines. It is desirable for the multilines to be radially far enoughapart to support the creation of interspaced interpolated lines, butclose enough together so that r.f. interpolation will not form artifactsdue to spatial undersampling. For example, if the steering delays arelimited to correspond to less than ±{fraction (λ/8)}, then the twosteered beams from each patch will fall within approximately the −1 dBwidth of the nominal patch beampattern. Also, because the steering delaybetween the left and right multiline on any element is thus limited to{fraction (λ/4)}, r.f. interpolated lines can be produced using a simpletwo tap interpolation filter ({fraction (λ/2)} delays would correspondto the Nyquist criterion). The {fraction (λ/8)} delay limitation limitsthe steering angle to approximately ±({fraction (λ/8)})/(4*λ) or{fraction (1/32)} radians. Thus the angle between the left and rightmultilines can be about {fraction (1/16)} radians, or about 3.6 degrees.If two other lines are symmetrically interpolated between the tworeceived multilines, the resulting line spacing is approximately 1.2degrees. A greater number of more closely spaced multilines orinterpolated lines can also be produced as desired.

[0073] In the 1D array example, instead of producing a single scanlinefrom each patch steered in the nominal steering direction, two scanlinesare produced, one steered slightly left of the nominal steeringdirection and one steered slightly right. In the case of a 2D array,several variations are possible. For a rectilinear 2D array, fourscanlines are produced for each patch, steered left, right, up and downin quadrature relationship. For a triangular-based 2D array such as ahexagonal array, three scanlines are produced at rotations of 120° asshown in FIG. 18d. The scanlines produced in this drawing are identifiedas B_(φ0), B_(φ120) and B_(φ240), respectively, where the subscriptnumber refers to the direction of rotation in the plane normal to thenominal steering direction of the patch and the angle φ is the smallangle at which each scanline is tilted from the nominal steeringdirection. The angle φ is kept small as described above so that thethree scanlines are kept within the beamwidth of the nominally steeredbeam. FIG. 18c illustrates a single scanline B₀ oriented normal to thepatch 202, as would be produced by the system shown in FIG. 18a, whichhas a beam nominally steered normal to the face of the patch 202.

[0074] Although the foregoing examples suggest the use of a rectangularscan geometry for a rectilinear array and a triangular scan geometry fora hexagonal array, the scan geometry is not intrinsically linked toarray geometry. A rectangular scan can be performed using a hexagonalarray and vice versa.

[0075] A system operating as illustrated by FIG. 18d is shown in FIG.18b. The scanhead in this drawing includes a 12 element patchmicro-beamformer which produces three multiline signals from each patch(B_(φ0), B_(φ120)and B_(φ240), for example) instead of one line as didthe micro-beamformer 280 of FIG. 18a. The micro-beamformed patchmultilines are sent over the n conductors of a cable 351 to theultrasound system's multiline beamformer 352. The multiline scanlinesfrom all of the patches are combined in the system multiline beamformer352 to form multiple scanlines. It is also possible to perform r.f.interpolation between the multiline scanlines. However, rather thancombine (beamform) the multiline signals from each patch and thenperform r.f. interpolation on the beamformed signals, it is preferredthat r.f. interpolation is performed on signals received from each patchseparately prior to beamforming combination. In this case, prior to theweighting and summation operations of r.f. interpolation, each patchsignal for each nominal steering direction is slightly delayed oradvanced by an amount determined by each patch position and the offsetof the interpolated line from the nominal line. The effect of the delaysis to maximize the coherence of the patch waveforms combined in the r.f.interpolation step. This reduces interpolation errors and improvessensitivity. Specifically, if N interpolated lines are produced from Mpatches, each patch having K multilines, then MN r.f. interpolators arerequired with each interpolator preceded by K delay states, one for eachmultiline. This same approach (i.e., delay+individual patch r.f.interpolation prior to patch signal combination) can also be used onpatch signals received from different directions in a non-multiline modeprovided that target motion between successive transmits is notexcessive. The multiple scanlines are then processed by the imageprocessor 284 and displayed on the display 100 as described previously.The number n of receive signal conductors of the cable is 768 if threemultilines from each of 256 patches are sent simultaneously withoutmultiplexing, a number which can be reduced by multiplexing if desired.The patch multilines received by the ultrasound system can beinterpolated to form additional scanlines prior to system beamformationif desired. However, since the processing of interpolation (weightingand summing) is mathematically compatible with that of beamformation,the patch multilines can be supplied directly to the system beamformerfor formation of beamformed multilines.

[0076] Several display formats may be used for the three dimensionaldisplay of the present invention. FIG. 20 shows a volumetric region 300which is being scanned by a 2D transducer array 200. The volumetricregion scanned can be in any desired shape, such as square, cylindrical,or pyramidal, depending upon the steering of the beams from thetransducer. In this example the volumetric region 300 is shown as ahexagonal pyramid. Shown within the volumetric region 300 is an imageplane 302, which is delineated by the double lines. The image plane 302is scanned in a time interleaved manner as the volumetric region 300 isscanned. The time interleaving enables the echo data from the imageplane 302 to be fully acquired in less time than that required to scanthe full volumetric region 300 and the frame rate of display of theimage plane 302 is thus greater than that of the volumetric display. Thetime interleaving of the volumetric and planar image data is illustratedby FIG. 21. This drawing shows a sequence E₃₀₀ during which echo data isacquired for the volumetric display. This sequence is periodicallyinterrupted during which echo data E₃₀₂ for the planar display isacquired. Some of the planar echo data can be used for both displays.The relative durations of the sequences and the number oftransmit-receive cycles needed for each display determine the frame raterelationship of the two displays.

[0077] The volumetric and the planar images are preferably displayedtogether as illustrated in FIG. 22. On the left side of the display 100is a three dimensional display of the volumetric region 300, which showsthe structure 304 in the volumetric region in a three dimensionalpresentation. On the right side of the display 100 is the twodimensional image plane 302, effectively showing a cut plane 306 throughthe three dimensional structure 304. While the frame rate of display ofthe three dimensional image 300 may be relatively low, the frame rate ofdisplay of the two dimensional image 302 will be much higher, which isuseful when diagnosing moving objects such as the heart. Preferably thelocation of the two dimensional plane 304 will be indicated in the threedimensional display, as shown in this example. This gives the user abasis of reference for the two dimensional image plane within thevolumetric region. The user has the ability to move the location of thecut plane 306 within the volumetric region so that a selected pathologycan be viewed at the higher frame rate. By manipulating a pointingdevice such as a mouse or trackball the position of the image plane 302within the volumetric region 300 can be changed on the left side of thedisplay. The user is given a choice of rotating the cut plane about thecenter axis of the volumetric region, or of dragging or moving the cutplane to an arbitrarily chosen position within the volume. Thus, thedisplay 100 displays a volumetric region at a relatively low frame rate,and a selected plane at a higher realtime frame rate. This methodapplies when the cut plane extends from the transducer aperture, thatis, the cut plane is not a “c” plane.

[0078] Another useful time interleaved display format is shown in FIG.23. Instead of interrupting scanning of the volumetric region 300 toscan an image plane, the scanning of the full volumetric region isinterrupted to scan a smaller volume 306 within the volumetric region300 for a higher frame rate of display of the smaller volume. Thescanning sequence is therefore E₃₀₀, E₃₀₆, E₃₀₀, E₃₀₆, E₃₀₀, and soforth. FIG. 23 shows the display of the full volumetric region 300.Outlined within the volumetric region 300 is the smaller volumetricregion 306. The smaller volume region is shown in an enlarged view atthe right side of the display. Since the smaller volume 306 is fullyscanned more frequently than the full volumetric region 300, the framerate of display of the smaller volume is greater than that of the fullvolumetric region. The number of beams and hence the time required toscan the smaller volume is a function of the lateral dimensions of thesmaller volume, the plane of FIG. 19, which in this example are thedimensions of top 308 and bottom areas of the smaller volume. Thus theframe rate of display of the smaller volume, and of both volumetricdisplays, can be increased by reducing the size of the top 308 of thesmaller volume. As in the previous example, the user may have the choiceof rotating the smaller volume about a center line within the volumetricregion or relocating the smaller volume to a desired arbitrary locationwithin the volumetric region 300. In this example the user haspositioned the smaller volume to encompass a portion of a coronaryartery 12 which is to be closely diagnosed for signs of obstruction,which may be more confidently done in the enlarged, higher frame ratesmaller volume image 306. Such a diagnosis would preferably be doneusing a Doppler mode, and preferably the power Doppler mode withsurrounding tissue rendered highly transparent or completely eliminated.

[0079] FIGS. 24-26 illustrate another display format which is useful forcoronary artery imaging as well as other vasculature. FIG. 24illustrates a volumetric region 300 which includes a three dimensionalimage of coronary arteries. A main artery 310 extends from the left sideof the volume and subdivides into branches 312 and 314. As shown anddescribed above, the coronary arteries follow twisting, tortuous pathsas they traverse the surface of the heart. A more confident diagnosiscould be obtained if these arteries could be effectively “straightenedout” for diagnosis. FIGS. 25 and 26 illustrate a technique for doing so.The clinician denotes a particular vessel for diagnosis, such as artery310. The ultrasound system them automatically traces the designatedvessel. One way to do so is illustrated in FIG. 25, in which theabscissa is the spatial dimension of the ultrasound image and theordinate is the intensity or brightness of the image. The curve 320illustrates the change in color or brightness across artery 310 from oneside of the vessel to the other. For example, the vessel may be coloredred against a gray background. The color red would increase as one sideof the vessel is encountered and the curve 320 rises at 310 a, anddecreases at the other side of the vessel at the downslope 310 b of thecurve 320. From slopes 310 a and 310 b the ultrasound system can readilydetermine the center 324 of the artery and can therefore trace along thecenter of the vessel in the image. If the automatic trace incorrectlybranches, such as following branch 312 when the clinician would like thetrace to follow branch 314, the clinician can click on branch 312 toerase its trace and click on branch 314 to continue the trace of artery310 onto branch 314.

[0080] Once the desired vessel path is identified, the vessel path isredisplayed in a straight path along its centerline 324 as shown in FIG.26. The vessel can be displayed in a cross-sectional view along thecenterline if desired or, since the vessel is in three dimensions in theimage of FIG. 24, the vessel can be “unwrapped” and the outercircumference displayed as the image height h in FIG. 26. When thevessel is shown in this “straightened” display and enlarged as desired,obstructions in the flow path such as that shown at 322 can be morereadily identified. Obstructions can often be more readily observed inan “unwrapped” display of the vessel circumference.

What is claimed is:
 1. An ultrasonic diagnostic imaging system forimaging three dimensional regions of a subject comprising: an ultrasonicscanhead including a two dimensional array of transducer elements whichacts to transmit ultrasonic waves and receive echo signals from avolumetric region of a subject; a multiline beamformer coupled toreceive echo signals from the elements of the two dimensional array toproduce beamformed echo signals; a scanline interpolator coupled to thebeamformer which acts to produce interpolated scanlines from thebeamformed echo signals; and a display coupled to the scanline processorwhich displays a three dimensional ultrasonic display.
 2. The ultrasonicdiagnostic imaging system of claim 1, wherein the scanline interpolatorcomprises an r.f. interpolator which performs interpolation of r.f. echosignals.
 3. The ultrasonic diagnostic imaging system of claim 1, whereinthe scanline interpolator acts to perform twopoint interpolation usingthe signals of two spatially distinct beams received from the volumetricregion.
 4. The ultrasonic diagnostic imaging system of claim 3, whereinthe scanline interpolator performs echo signal interpolation utilizingweighting factors which are a factor of two (2⁻¹).
 5. The ultrasonicdiagnostic imaging system of claim 4, wherein the weighting factors arein the proportions of ½:½.
 6. The ultrasonic diagnostic imaging systemof claim 4, wherein the weighting factors are in the proportions of ¼:¾.7. The ultrasonic diagnostic imaging system of claim 6, wherein thescanline interpolator acts to perform two levels of interpolation; andwherein the weighting factors used in the second level of interpolationare in the proportions of ¼:¾.
 8. The ultrasonic diagnostic imagingsystem of claim 1, further comprising a filter, coupled to the scanlineinterpolator, which acts to filter interpolated scanline data.
 9. Theultrasonic diagnostic imaging system of claim 1, wherein the scanlineinterpolator comprises a linear interpolation filter.
 10. The ultrasonicdiagnostic imaging system of claim 1, wherein the scanline interpolatorutilizes an interpolation kernel that has a non-linear shape.
 11. Theultrasonic diagnostic imaging system of claim 1, wherein the multilinebeamformer produces multiple beams from the same transmit event; whereinthe scanline interpolator acts to produce interpolated scanlines frommultiline echo signals.
 12. The ultrasonic diagnostic imaging system ofclaim 1, wherein the multiline beamformer is partitioned between thescanhead and an ultrasound system processor.
 13. A method forultrasonically imaging a volumetric region of a subject comprising:transmitting ultrasonic energy into the volumetric region; receivingecho signals in response to the transmitted ultrasonic energy by theelements of a two dimensional array transducer; beamforming the receivedecho signals to produce spatially distinct beams of the volumetricregion; interpolating the spatially distinct beams to increase thescanline density of the image data of the volumetric region; anddisplaying a three dimensional ultrasonic image of the volumetric regionwhich utilizes interpolated scanline data.
 14. The method of claim 13,wherein interpolating further comprises performing two-pointinterpolation of two spatially distinct beams.
 15. The method of claim13, wherein interpolating further comprises weighting the echo signalsof spatially distinct beams by weighting factors which are a factor oftwo (2⁻¹).
 16. The method of claim 13, further comprising filteringinterpolated scanline data.
 17. The method of claim 13, whereininterpolating further comprises performing r.f. interpolation of echosignal information.
 18. The method of claim 13, wherein interpolatingfurther comprises performing linear interpolation.
 19. The method ofclaim 13, wherein interpolating further comprises utilizing aninterpolation kernel that has a non-linear shape.
 20. The method ofclaim 13, wherein the array transducer is located in a scanhead andwherein beamforming comprises forming partially beamformed signals inthe scanhead and performing further beamforming of the partiallybeamformed signals in an ultrasound processor coupled to the scanhead,wherein interpolating is performed in the ultrasound processor.